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We discuss the development of dynamic 3D bioreactor-based
systems as in vitro models for use in DMPK studies.
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The use of bioreactors as in vitro models in
pharmaceutical research
Maaria Ginai1, Robert Elsby2, Christopher J. Hewitt1,
Dominic Surry2, Katherine Fenner2 and Karen Coopman1
1
Centre for Biological Engineering, Department of Chemical Engineering, Loughborough University,
Loughborough LE11 3TU, UK
2
In vitro In Silico ADME, Global DMPK, AstraZeneca Research and Development, Alderley Park, Mereside,
Macclesfield SK10 4TG, UK
Bringing a new drug to market is costly in terms of capital and time
investments, and any development issues encountered during late-stage
clinical trials can often be the result of in vitro–in vivo extrapolations
(IVIVE) not accurately reflecting clinical outcome. In the discipline of drug
metabolism and pharmacokinetics (DMPK), current in vitro cellular
methods do not provide the 3D structure and function of organs found in
vivo; therefore, new dynamic methods need to be established to aid
improvement of IVIVE. In this review, we highlight the importance of
model progression into dynamic systems for use within drug development,
focusing on devices developed currently in the areas of the liver and blood–
brain barrier (BBB), and the potential to develop models for other organ
systems, such as the kidney.
Currently, the pharmaceutical industry balances the stringent testing of drugs and products
against soaring costs and return on investments risks. Bringing a new drug to market will,
presently, on average cost US$1.3 billion over 12 years [1,2] with preclinical in vitro testing
consuming approximately half of the total development time [2]. Moreover, many new
chemical and biological entities (NCEs and NBEs, respectively) that fail late-stage human testing
provide evidence for the fact that pharmacological and toxicity data from in vitro cell-based
assays are not always predictive of the clinical situation [3]. Owing to these financial and time
commitments, there is great emphasis within the industry on the development of newer and
more reliable in vitro and preclinical methods to accompany or replace the existing methods of
NCE–NBE investigations.
In vitro models, particularly cell-based versions, are used in various areas of drug discovery, such
as target identification and validation using disease models, compound screening and basic
cytotoxicity using static cultures, through to absorption, distribution, metabolism, excretion and
toxicology (ADMET) studies on lead compounds. Although high-throughput cell-based assays
have revolutionised the efficiency and speed at which compounds can be screened, ADMET
models ideally should be an accurate representation of the physiological or pathophysiological
Maaria Ginai is currently
studying a BBSRC CASEfunded PhD at the Centre for
Biological Engineering at
Loughborough University
(UK) in the area of bioartificial device progression to
in vitro models for use in the
pharmaceutical industry. Her
PhD is supported by AstraZeneca. She graduated with
a first-class Bachelors degree with honours in biomedical science from the University of Kent (UK) in
2010.
Christopher J. Hewitt is
the director of the £7.3 M
EPSRC Doctoral Training
Centre in Regenerative
Medicine and co-founder of
the £2 M Centre for Biological Engineering at
Loughborough University
(UK). He also leads the Cell
Technologies research
group, whose work spans the engineering–life science
interface seeking to understand the interaction of the
organism with the engineering environment within
such diverse areas as microbial fermentation, biotransformation, cell culture and, mostly recently,
regenerative medicine bioprocessing. Christopher has
a first-Class Bachelors degree in biology from Royal
Holloway College, University of London (UK) and a
PhD in chemical engineering from the University of
Birmingham (UK).
Karen Coopman was
appointed to a lectureship at
Loughborough University,
where she is the operations
manager of the EPSRCfunded Doctoral Training
Centre in Regenerative
Medicine. She co-leads the
Cell Technologies research
group within the Centre for
Biological Engineering of the University, and is currently a member of the Early Career Forum in
Manufacturing Research of the EPSRC. The overarching themes of Karen’s research are the manufacture of cellular therapies and the use of cells in the
drug discovery process. Karen has a first-class
Bachelors degree in pharmacology from the University of Bristol (UK) and a PhD from the Department of
Pharmacy and Pharmacology at the University of Bath
(UK).
Corresponding author:. Coopman, K. (k.coopman@lboro.ac.uk)
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Drug Discovery Today Volume 18, Numbers 19/20 October 2013
Limitations of current 2D systems
In vitro cell-based models offer several advantages over most cellfree methods because they do not require purification of the target
protein, they immediately select against compounds that are
generally cytotoxic or cannot permeate cell membranes, and
enable measurement of functional transport activity that is more
representative of the physiological state [5]. Using primary human
cells in these assays is the ideal gold standard [6]; however, freshly
isolated cells, cell cultures or tissue fragments cannot be used to
study effects that occur over a longer time period owing to the loss
of functional integrity from drug transporter loss and enzyme
depletion after a few hours [7,8]. In addition, primary human cells
are a scarce and precious resource. For these reasons, immortalised
cell lines, such as (transfected) MDCKII-MDR1 or Caco-2, are often
used in favour of primary cells. Such cells are relatively inexpensive
and easy to propagate, especially in high-throughput systems;
however, owing to the genetic transformation process, aspects
of the original cell that are important to ADMET studies might
be lost, such as transporter expression and function, and cell
adhesion molecules [8].
Although relatively simple and effective when using the appropriate cell type, 2D in vitro cell-based assays encounter the traditional problems of static culture: a constantly changing
environment owing to nutrient depletion and metabolic waste
accumulation. Strategies to improve such systems have been
introduced based on mimicry of the natural physiological environment of the organ in question. The re-establishment of heterotypic cell–cell and cell–extracellular matrix interactions through
co-cultures and sandwich cultures of cells, respectively, have
proven useful in the progression of in vitro model development
for several tissue types, including liver [9,10], brain [11] and lung
[12]. These models show improved functionality and growth
compared with their single culture counterparts, whilst keeping
their ease of propagation and use compared with primary cell
cultures, which is imperative when developing long-term in vitro
models for pharmaceutical applications.
Perfusion systems of isolated organs or tissue fragments offer the
closest ex vivo model of the natural physiological state, alleviating
the problems encountered with static culture, maintaining 3D
architecture and prolonging the functionality, heterogeneity,
structural complexity and cell–cell interactions [13]. However,
the shortage of organs available and the costs and difficulties of
maintaining primary cells ex vivo make perfusion cultures unsuitable as a routine in vitro cell-based system.
Moving towards 3D systems
The development of 3D constructs for pharmaceutical research
now encompasses a variety of systems and devices. 3D cell culture
with cell types native to a specific organ that are grown on natural
scaffolds, such as collagen, decellularised scaffolds or synthetic
biodegradable scaffolds, present a bridge between 2D culture and
the in vivo cellular environment. 3D devices for pharmaceutical
research, with structural complexity and native functionality, are
becoming available; however, these cultures are difficult to maintain in vitro and still lack the vascular component of tissues.
Extracellular matrix (ECM) coatings that release growth factors
and that can be coated onto scaffolds (e.g. polymer scaffolds) have
been investigated to aid cell attachment, growth and partially
represent the native environments, which can have profound
effects on the morphology, behaviour and functions of cells
needed in drug discovery processes; however, further work is
needed to progress these cultures for use as models [14].
Bioartificial devices combine flow-through systems, usually a
form of simple plug flow bioreactor, with cells, thereby mimicking
both the dynamic and cell-specific aspects of native tissues [13].
They are being incorporated in the pharmaceutical industry owing
to the potential payoff of getting more drugs to market.3 As we
discuss below, these devices are being developed from clinically
used systems and, although they offer cellular and environmental
similarities, most systems currently lack the complex architecture
and microenvironments seen in 3D constructs.
Status of bioreactors in the development of in vitro
models
Bioreactors are used in a range of applications, from the production of biopharmaceuticals to applications in tissue engineering,
such as cell expansion, generation of 3D tissue constructs and
direct organ support devices [15] (Table 1). Unlike 2D static
cultures, bioreactors provide a controllable environment in terms
of pH, temperature, nutrient supply and shear stress for any cells or
cellular constructs incorporated into them. However, key limitations in 3D culture, such as mass transfer of oxygen and nutrients
to cells, still apply. For example, cardiac muscle tissue contains a
3
Cuddihy, M. (2011) 3D Cell Cultures Prevalent in Pharma: Human Predictive
Functional Tissue Models Conference. Available from: http://3dbiomatrix.
com/2011/12/01/3d-cell-cultures-prevalent-in-pharma-human-predictivefunctional-tissue-models-conference/ [Accessed 21 May 2013].
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state, which might be lacking in 2D static cultures. Tissue engineering advances show promise in the development of models
with organised structure and cell types similar to in vivo. However
the dynamic aspect of in vitro models is yet to fully be addressed,
although it is starting to be investigated with the use of 3D
bioreactors seeded with cells.
Within the pharmaceutical industry, the absorption, distribution, metabolism and excretion properties of NCEs are typically
investigated as part of DMPK science. These DMPK properties are
optimised within drug discovery so that an NCE has the right
clinical profile; for example, ideally, high oral bioavailability, an
elimination half-life that enables a once-daily dosing regimen,
sufficient exposure at the target tissue and an absence of drug–drug
interaction (DDI) potential [4]. In vitro 2D cell-based models are
routinely used within pharmaceutical research to investigate the
DMPK properties of absorption (Caco-2, derived from human
colon carcinoma cells), distribution (MDCKII-MDR1, derived from
canine kidney cells) and metabolism and/or excretion (hepatocytes) for NCEs. More recently, some 3D models have started to be
investigated to assess their utility in studying aspects of ADMET,
namely the bioartificial liver (distribution, metabolism and elimination), the BBB model (distribution) and the bioartificial kidney
(excretion). Therefore, in this review, we focus on the potential of
bioartificial devices that have been developed for therapeutic use
to progress to pharmaceutically relevant in vitro models for use in
DMPK studies.
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TABLE 1
Bioreactor designs and uses in tissue-engineering applications
Bioreactor type
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Pros
Cons
Uses
Refs
Induces movement of oxygen and
nutrient within culture medium
Easy to scale up
Perception that shear stress can
have damaging effects on cells
Poor mass transfer when applying
to static scaffolds
Growth of cartilage constructs
[18]
Efficient surface area for attachment
Potential for immunoisolation
Perception that cells are
protected from shear
Nonuniform cell distribution
Fibre membrane can act as a
physical transport barrier
Generation of a bioartificial liver
support system
[19]
Uniform cell distribution
Good distribution of nutrients
and gas transfer
Stable microenvironment
Complex to scale up
Cells exposed to shear stresses
Low surface area:volume ratio
Cartilage construct engineering
from chondrocyte-seeded scaffolds
[20]
Promotes 3D structure formation
Reduced physical transport barriers
Easy to scale up
Nonuniform distribution of cells
Cells exposed to shear stresses
Nutrient and oxygen transfer issues
Early bioartificial liver system
studies
[21]
high density of cells and has a high oxygen demand, which in vivo
is supplied by a dense network of capillaries through the tissue. To
gain a better understanding of the native environment of cardiac
tissue and its influence on the synchronously contracting cardiac
constructs, a bioreactor-based model has been developed [16].
Incorporated into this device are rodent cardiomyocytes and
fibroblasts grown on scaffolds with a parallel array of perfused
channels to mimic the capillary network. Blood is mimicked by
supplemented media and the model can be used to assess several
parameters, such as the impact of perfusion rates of oxygen on
viable cell density, to define the conditions needed to culture
cardiac constructs with clinically relevant thickness [17].
Hollow fibre and perfusion systems in particular are widely used
in the generation of in vitro models. In addition to enabling users to
investigate the impact of environmental conditions on cells, they
can also be used to model pharmacological interactions of drugs
with cells in infectious [18] or model disease [19] states.
Progression from these systems has generated multicompartmental bioreactors that mimic the physiological environment
more closely. As well as being used clinically [20], they are being
utilised as in vitro models, such as the scaling down of a 3D
perfusion multicompartment hollow fibre liver bioreactor for
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use in in vitro pharmacological studies [21]. This bioreactor comprises three interwoven sets of hollow fibre, two for countercurrent medium perfusion and one for gas supply (Fig. 1). Primary
human liver cells (both parenchymal and nonparenchymal) are
co-cultured in the extra capillary space (ECS), and have been
shown to exhibit prolonged cytochrome P450 (CYP) enzyme
activity when maintained for over 3 weeks, enabling long-term
pharmacological studies to be carried out. Hollow fibre bioreactor
(HFB)-based models are also being developed for the BBB and
kidney. It is important to assess critically the individual features
of established systems so that optimised models for pharmaceutical research can be generated, and this is the focus of the
remainder of the review.
Perspectives from established systems used within
pharmaceutical research
Bioartificial liver devices
The liver is the main excretory organ within the body, but as well
as eliminating waste products through the bile, it also carries out a
variety of highly specialised functions, including oxidative detoxification, intermediate metabolism and supply of nutrients, modulation of immune and hormonal systems, and protein and
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(a)
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(c)
2 ml 8 ml 800 ml
Drug Discovery Today
FIG. 1
(a) Smallest capillary unit with medium capillaries that are independently perfused (red and blue) and one gas capillary (yellow); cells are cultured within the extracapillary space (cell compartment). (b) The bioreactor comprises three interwoven capillary bundles, each made of multiple hollow fibre capillaries for countercurrent medium perfusion (red and blue) and gas supply (yellow), which enables decentralised nutrient and O2–CO2 exchange with low gradients. (c)
Downscaling of the clinical-scale bioreactor prototype with a cell compartment volume of 800 ml was realised by reducing the length and number of capillary
layers in two axes of 3D space within the bioreactor, resulting in a laboratory-scale bioreactor with a cell compartment volume of 8 ml and a further downscaled
model with a cell compartment volume of 2 ml.
Reproduced from [21].
macromolecule synthesis [6]. Disruption in these functions owing
to liver disorders, such as acute or chronic liver failure, can lead to a
possible 90% mortality rate if essential hepatic functions are not
restored during the crucial phase of liver failure [22]. Therefore,
there is great emphasis on the development of reliable in vitro
models to assess drug interactions and pharmacological responses
to prevent drug-induced acute liver toxicity, and to model liver
disease states accurately for the development of NCEs targeting
disease-induced liver failure.
Bioartificial liver (BAL) devices have gone through many iterations while being developed as an extracorporeal device for the
treatment of liver failure [23]. Early flat plate and monolayer
bioreactors exhibited uniform cell distributions and microenvironments, but presented complex scale-up issues and a low surface
area:volume ratio [24,25]. Perfused beds and/or scaffolds promote
3D architecture of seeded cells, which when considering the
complex architecture of the liver, is an advantage. The cells
showed improved metabolic and synthetic function compared
with cells cultured in a spinner flask [26]; however, exposure to
shear forces and nonuniform perfusion of medium require an
improved bioreactor design. Conventional hollow fibre cartridges
offer increased surface area for attachment and immunoisolation
of cells when seeded in the ECS [27]. Cells are also protected from
shear as medium flows through the fibres. However, nonuniform
cell distribution along the fibres is a common problem and,
although being beneficial in terms of support and improved cell
growth and performance if adapted to the specific cell type [28],
membranes can present a physical barrier against the transport of
nutrients and metabolites. As mentioned previously, development
from standard HFBs to multicompartmental devices, such as that
in Fig. 1, has been a natural progression, given the native architecture of the liver, and has been utilised for both clinical and
pharmacological applications [21].
Cell sources
Hepatocytes are the primary choice for use in BAL systems
because they are recognised to be the closest cellular model to
the liver [29,30], and carry out most of the in vivo metabolic
processes, including the synthesis and excretion of albumin,
metabolism of amino acids, urea production and the processing
and elimination of drugs and toxins. Phase I and phase II
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drug-metabolising enzymes are crucial in the processing of drugs
within the liver; therefore, preservation of these enzymes within a
model is essential to reproduce cellular environments in vivo.
Cell lines have frequently been utilised for in vitro studies owing
to prolonged cell survival and ease of maintenance compared with
primary hepatocytes. Popular liver-derived cell lines, such as the
HepG2/C3A cell line derived from hepatomas, have been used in
pharmacological studies, and have recently been utilised in microfluidic bioreactors to examine the sensitivity of the models in
response to various xenotoxins [31,32]. However, compared with
primary hepatocytes, functionality is often lost or altered in these
cell lines [6]. For instance, phase I enzymes were expressed at
significantly lower levels in HepG2 cells relative to primary human
hepatocytes [33]. Therefore, primary cells would be ideal for
incorporation into a pharmacologically relevant in vitro model.
However, recent work has demonstrated the successful incorporation of a HepaRG human hepatoma cell line into a 3D BAL device,
in which metabolic and transporter functionalities of the cells are
preserved from a mere few hours in suspension to at least 1 week in
the 3D device, thereby enabling performance of longer-term in
vitro pharmacological studies [34–36].
Primary porcine hepatocytes have been widely incorporated
into BAL devices [37,38] because they are readily available, are
able to maintain differentiated metabolic functions under certain
culture conditions, provide a high yield of cells [39] and have a
similar metabolic profile to human hepatocytes [40]. However, the
environmental triggers that sustain differentiated hepatic function in vitro have not yet been fully characterised [41]. Utilising
porcine cells addresses the problem of finding a sustainable source
of hepatocytes; however, it would be important to obtain cells
derived from a human source to establish a representative model of
the natural physiological environment for in vitro pharmacological
studies in both normal and diseased states.
Primary human adult hepatocytes are considered to be the gold
standard for biocompatibility and functionality within BAL systems [6,42], but exhibit a rapid, time-dependent, general loss of
function in vitro [7], including reduced cytochrome P450 and
other biotransformation enzyme activities [43] and the downregulation of liver-enriched transcription factors, such as CCAAT/
enhancer binding protein alpha [42]. However, the use of complex chemically defined media has enabled hepatocytes to go
through several cycles in culture, as well as the stabilisation of
hepatocyte morphology, survival and liver-specific functions
[37,42,44,45]. Co-culture with nonparenchymal cells has also
been shown to preserve liver-specific functions, including initiating physiological reorganisation of cells and the distribution of
canalicular transporters similar to the pattern found in vivo [21].
The mechanisms behind these changes are not fully understood,
but highly conserved signalling pathways are thought to be
involved [9].
There is also limited availability of adult human hepatocytes
owing to their acquisition from liver biopsies, which has presented
problems with utilising this source in clinically used devices
because the number of cells needed for seeding a BAL for the
treatment of hepatic failure is large, approximately 2 1010
[25,46]. However, the number of cells needed for bioreactor-based
models is approximately 3 108 per device [21], which could be
attained with minimal culture in vitro.
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Device designs and functionality
Within the native liver, hepatocytes are ordered in a 3D lobular
network of hexagonal constructs, surrounding a central blood
vessel. Individual cells have polarity between the apical domain
interfacing with bile canaliculi and a basal domain interfacing
with the sinusoid, helping localise specific functions. During
culture, hepatocytes flatten and spread upon attachment, whether
grown on a flat surface or a scaffold. Although cell–cell junctions
can be formed, the cytoskeleton is largely disrupted, and leads to
loss of cell polarity and, ultimately, liver-specific functions [6].
Nonetheless, C3A cells and primary porcine cells have been
utilised within HFBs in clinically tested devices. The extracorporeal liver assist device (ELAD) comprises a dual pump dialysis
system connected to one or more hollow fibre cartridges seeded
with C3A cells in the ECS. Blood is ultrafiltered and pumped
through fibres, with plasma flowing through the ECS. This direct
interaction enables cellular processes to occur before the ultrafiltrate is filtered by a dual membrane to remove cells and cellular
debris and returned back into the blood stream [47].
The liver has an oxygen-rich environment, fed by the hepatic
arteries and portal vein. Unsurprisingly, an improvement in functionality of C3A cells in HFBs was noted when red blood cells were
added in culture medium [48], highlighting that an adequate
oxygen supply is a prerequisite for the maintenance of hepatocytes
in vitro. Considering this, the Excorp medical bioartificial liver
support system (BLSS) incorporates an oxygenator and heat
exchanger along with a blood pump and HFB seeded with primary
porcine hepatocytes in the ECS [49]. Whole blood flows through
the fibre, protecting the cells from host immunological responses
[38].
Another BAL, the modular extracorporeal liver support system
(MELS), has attempted to take into account the biliary system of
the liver, which excretes conjugated bilirubin and metabolites.
Hollow fibre bundles, seeded with human hepatocytes from discarded donor livers, are incorporated alongside a detoxification
module for removing albumin bound toxins. This provides an
inlet of plasma and contact with hepatocytes in the ECS, an outlet
of plasma, and hydrophobic membrane bundles for gas exchange
inside the bioreactor [50]. The structure of the seeded bioreactor
module is an early iteration of the multicompartmental bioreactor
being developed for in vitro pharmacological studies by Gerlach
et al. [6] (described above and in Fig. 1).
These systems, although showing positive biochemical changes
in patients and maintaining patients through to transplantation,
show no significant effect on survival and, therefore, can be
assumed to be unsuitable replicates of the environment in vivo.
Aspects such as co-culture of nonparenchymal and parenchymal
cells, efficient oxygenation of cells and the presence of a biliary
system for the removal of toxic protein-bound substances and
other metabolites should be taken forward in the generation of
both efficient in vitro models, as has been demonstrated in the
multicompartmental liver bioreactor (Fig. 1), and clinically relevant medical devices.
Use of BALs in DMPK studies
Bioreactor-based models are currently being integrated into the
DMPK process, initially assessing the functionality of cells within
the systems relative to drug transporter and metabolic enzyme
expression and function. When investigating the metabolism of
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For both the medical devices and in vitro models highlighted
above, there are currently no standardised criteria to define the
efficiency and efficacy of these systems. Furthermore, any criteria
applicable to a device will be partly dependent on the type of
bioreactor used (i.e. hollow fibre systems require tight monolayers
of cells on the extra luminal side of the fibre membrane). However,
cellular metabolic functions and enzymatic and transporter retention are paramount in all bioartificial liver systems. From the
clinically relevant devices mentioned above, there is little focus
on the performance of cells, and factors used to determine the
efficiency of the systems before being used in clinical studies are
not clearly stated. Devices that have been utilised for DMPK
research go further in defining the criteria used to assess the
models, such as monitoring metabolic activity via albumin synthesis, lactate dehydrogenase or aspartate aminotransferase, and
drug-metabolising enzymatic activity via administration of drugs
with cytochrome P450-dependent metabolism, as well as drug
transporter functionality and visualisation by inhibitor transport
assays and staining [21,34–36,51–53]. With the progression of
these systems, both medical and in vitro models, standardised
criteria defining the performance relative to native conditions
of the above-mentioned factors should be produced.
BBB models
The BBB is one of the most important lines of defence against
infections in the brain, acting as a physical and metabolic barrier
between the central nervous system (CNS) and systemic circulation and maintaining homeostasis within the brain. The capillary
structure is distinct from that found in peripheral tissues, with the
presence of tight junction proteins between endothelial cells
(ECs), the restriction of transcellular movement of molecules
and contact of ECs with astrocytes to provide support [54], protection from hypoxia and aglycaemia and to separate capillaries from
neurons [55].
The decreased paracellular permeability of the BBB, while having multidrug transporters such as P-glycoprotein (P-gp, also
known as MDR1), prohibits the transport of large drug molecules
into the CNS, presenting the need for a reliable in vitro model for
ADME studies. Existing models have not been characterised as well
as liver devices, partly owing to their lack of clinical relevance
regarding the generation of an extracorporeal device, but also
owing to the complexity of cell alignments and functionality
within the model. However, similar to the BAL devices, HFBs have
been developed and utilised as ‘dynamic’ models of the BBB [56].
Model designs
Unlike liver models and bioartificial devices, there is no one cell
type that can be incorporated to mimic the BBB. Most models
reported in the literature use two cell types; ECs and glial cells
(astrocytes). Early models simply used mono- and/or co-cultures of
ECs alone or with glial cells cultured on transwell plates. This
approach is attractive owing to its simplicity; however, primary
ECs lose their BBB properties in vitro owing to the absence of
stimuli that are normally present in vivo [56]. Furthermore, specific
transporters, such as P-glycoprotein (P-gp), are downregulated in
the absence of astrocyte-derived soluble factors, leading to abnormal permeability across the monolayer [57]. This dedifferentiating
process might also be accelerated by nonphysiological conditions,
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diclofenac, freshly isolated human hepatocytes were found to
produce 90% less hydroxydiclofenac, the hydroxylated product
of diclofenac generated directly by the cytochrome P450 enzyme
CYP2C9, after a 3-day culture in 2D conditions compared with a
high level of metabolites from human hepatocytes in the multicompartmental bioreactor described above [21,34]. The major
intermediary diclofenac metabolite, diclofenac acyl glucuronide
[produced from the glucuronidation of diclofenac by UDP-glucuronosyltransferase (UGT) enzymes before hydroxylation by
the P450 enzyme CYP2C8], was found to be absent in human
hepatocyte suspensions, but present in the hepatocyte-seeded
bioreactor, possibly owing to the bioreactor system mimicking
the in vivo environment and enabling the efflux of metabolites
away from the cells, thereby avoiding further metabolism. Overall,
the relative levels of the metabolites found in the suspensions versus
the 3D cultures were similar; however, all human in vivo metabolites
were detected in the bioreactor culture, whereas one major in vivo
metabolite was not present in the hepatocyte suspension. In addition, the biotransformation capacity of the hepatocytes within the
bioreactor was retained for at least 1 week, providing a good model
for metabolite investigations from slowly metabolised drugs [34].
The activity of CYP3A4 was demonstrated in the 3D bioreactor
system by monitoring the metabolism of atorvastatin acid (ATA)
to its 2- and 4-hyroxylated metabolites [35], and for HepaRG cells in
the 3D system by monitoring midazolam, which indicated that
CYP3A4 activity was retained for up to 2 weeks [47]. This retention of
CYP3A4 activity is longer than that currently available in other in
vitro models.
Transporter expression is also a key aspect in the functionality of
hepatocytes for DMPK screens. Within the bioreactor, immunohistochemical staining of multidrug resistance-associated protein
(MRP2) and multidrug resistance protein (MDR1), both bile canicular efflux transporters, showed expression, localisation and distribution similar to native liver tissue for at least 2 weeks within the
bioreactor [21,36]. Expression of the basolateral uptake transporter
organic anion transporting polypeptide 1B1 (OATP1B1) has also
been investigated on hepatocytes within the 3D system preserved
for 9 days, functional activity of OATP1B1 was demonstrated by
the transport of substrates estradiol- 17b-D- glucuronide (E17bG)
and atorvastatin acid (ATA) in the presence and absence of the
inhibitor estrone-3-sulphate (E3S), and was retained for 7 days
compared with the decrease of functional activity of OATP1B1/
1B3 in plated primary human hepatocytes after just 2 h [35,51].
When assessing drug–drug interactions, OATP1B1 is a vital hepatic
transporter to evaluate, because it is inhibited by many commonly
used drugs, and is responsible for the targeted uptake of the widely
prescribed statins to their site of action [52]. This has been demonstrated by similar increases in the elimination half-life of ATA in
the presence and absence of OATP1B1 inhibitors, with a 2.7-fold
increase in the presence of rifampin [53], and a 1.7-fold increase in
the presence of E3S within the bioreactor model. The difference in
these values had been attributed to experimental setup and potencies of the two inhibitors, and it is also been noted that the major
limitation of this model lies in the variable quality of isolated
human hepatocytes from different donors [35]. However, these
initial studies into the functionality of drug transporters and
metabolic enzymes clearly establish the potential of this model
for use throughout the DMPK process.
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TABLE 2
Performance of various cell types within the HFB in the generation of an in vitro BBB model
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EC
Glial cell
Days prolonged in culture
Refs
Rat brain microvascular ECs
Rat brain astrocytes
30
[62]
Human brain ECs
Human foetal astrocytes
Not reported
[62]
Bovine aortic ECs
C6 glial cell line
30
[62]
such as the presence of serum on both sides of the polarised EC
monolayer instead of only on the luminal surface in vivo. Coculture models incorporating glial cells mimic the in vivo situation
more closely, demonstrated by the increased expression of brain
endothelial marker enzymes, transporters, and an increased transendothelial electrical resistance (TEER) across the barrier indicative of a ‘tight’ monolayer. This model is effective for studies
around the function of the BBB and processes and interactions
between the ECs and the glia [56]. However, for a replicative model
of the BBB, factors such as shear stress, which is vital to promote
growth inhibition and differentiation of ECs and enables the
trafficking of metabolic fuels to the brain, should be incorporated
into animal-based models before ideally progressing into humanbased models, to enable direct extrapolation of data from studies,
and functional and physiological mimicry of the human BBB in
vivo [58].
Dynamic (flow-based) models incorporate HFBs with coated
fibres seeded with ECs (intraluminally) and glial cells (extraluminally). In an early model developed by Stanness et al. [59], ECs
and corresponding glial cells from human, rat and bovine sources
were cultured successfully in a HFB coated with ProNectin (Table
2) [59]. Morphology and growth patterns were consistent with all
cell types; however, it was noted that when differing types of EC
and glial cells were cultured, such as human umbilical cord ECs
with C6 cell line glial cells, BBB properties were not developed.
Poor EC growth in the absence of glial cells within the hollow
fibres was observed via low transendothelial resistances and studies assessing the transport of morphine demonstrated permeability ratios to be similar to in vivo, but absolute values to be
lower. This was thought to be because of the differences in flow
rate patterns in the devices and from in situ. The results supported
the use of HFBs in BBB models; however, because bovine aortic
ECs + C6 seeded cartridges were used for drug transport studies,
utilising human brain ECs in the HFBs would be a natural progression.
Neuhaus et al. [60] utilised a commercially available bioreactor
from the same manufacturer as Stanness et al., and seeded immortalised porcine brain microvascular endothelial cells (PBMEC/C12) intraluminally with the glial cell line C6 cells in the extra
capillary space. Medium was pumped through the system applying
shear stresses of 2.7–3.9 dyn/m2, mimicking conditions in brain
capillaries. Results of cell growth and transport of benzodiazepines
were compared between the transwell and flow-based models. The
results showed that PBMEC/C1-2 cells seeded on the transwell
inserts started detaching after 5 days, prohibiting the use of
transwell systems in long-term experiments, whereas growth
and morphology of the ECs in the HFB was affected owing to
exposure to flow, prolonging the survival of the cells within the
device. Permeability coefficients of benzodiazepines were also
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shown to be lower in the HFB model, presumably owing to tighter
monolayers and improved barrier functions within the device [60].
The BBB is an active site in terms of transport, expressing
multiple drug transporters to restrict compounds with high affinity for efflux transporters and low transcellular membrane permeability into the CNS. However, owing to its complexity, there
are few established in vitro models, and CNS-targeted compound
analysis relies heavily on in vivo and in silico assessments. P-gp
(MDR1) is a predominant efflux transporter in the BBB, capable of
limiting the penetration of a wide range of substrates into the
brain. For this reason, MDCK–MDR1 cellular bidirectional transport assays are currently used in DMPK screening to assess both
CNS-penetrating and CNS-restricted compounds, as appropriate
depending upon target profile [4]. Differences in transporter
expression and passive permeability have been shown in 2D static
co-culture models (described above). However, investigations have
also highlighted interspecies variability, with a twofold higher
permeability in a rat EC and glial model than in the human
counterpart. The limitations of the current transwell model utilising transfected cell lines are its static nature and lack of availability
of human cells. However, this system could be improved by using
rat ECs and glial cells, to better represent the in vivo environment
than is currently available with transfected cell lines alone, to
account for the difference in permeability [61]. Owing to the
nature of individual transporter testing in current 2D cultures,
3D models exhibiting BBB characteristics could also be utilised for
drug–drug interaction studies as the 3D liver models are starting to
investigate.
Overall, the outcomes of these models are promising for the use
of HFBs in pharmacological applications. However, the human
brain is more complex than its rodent counterpart and, although
many drug transporter proteins are expressed in rodent ECs [56],
primary human ECs and glial cells would ideally be incorporated
into an optimised device. This might be difficult to generate for
large-scale studies because healthy human glial and ECs are rarely
available, so one possible source would be the differentiation of
stem cells into these cell types. However, extensive work in stem
cell differentiation and characterisation from embryonic and
induced pluripotent stem cells is needed, and is not a viable
source for models at present. Recreating aspects of the microenvironment of the CNS in terms of ECM proteins surrounding glial
cells might also improve performance of the cells within the
device.
Progressions in bioartificial kidney (BAK) devices as an
in vitro model
The need for an optimised model from existing clinical devices
The kidneys are primarily responsible for the excretion of waste
products, such as urea and ammonia, from the body through the
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(b)
200ml/min
V
A
A
V
Continuous
haemofilter
150ml/min
10ml/min
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4ml/min
Waste
185ml/min
7ml/min
10ml/min
Haemofiltration
cartridge
Bioartificial
cartridge
Bioartificial
cartridge
(RAD)
3ml/min
5ml/min
Waste
Waste
Drug Discovery Today
FIG. 2
Schematic diagram of (a) bioartificial kidney device used for the treatment of acute renal failure developed by Humes et al. [66]. (b) Bioartificial kidney device
developed as a renal replacement therapy for chronic renal failure by Saito et al. [65]. The devices have similar components and configurations; however the
device developed by Saito et al. incorporates a continuous haemofilter for continuous treatment.
Based on device layouts from [65,66].
urine. They also provide a variety of significant secondary functions
through homeostatic, endocrine, immune and metabolic processes.
Chronic kidney disease (CKD) is a growing problem within the
UK, with approximately 1.8 million people with CKD in England
alone. This is putting a growing burden on the NHS, costing over
£1.4 billion per year, which is largely because of renal replacement
therapy (e.g. dialysis), which showed a 29% increase in usage from
2002 to 2008, and complications, such as heart disease and stroke
[62]. This highlights the need for early diagnosis of the disease,
more efficient and cost-effective renal replacement therapies,
more effective drug treatments and, consequently, an optimised
kidney model for pharmacological studies.
Existing bioartificial kidney (BAK) devices have been developed
primarily as a renal replacement therapy. They incorporate both
mechanical and biological processes, pumping haemofiltered
blood through a renal assist device (RAD). Coined by Humes
et al., this device contains renal cells seeded in the lumen of
hollow fibres of a standard haemofiltration cartridge [63], contrary
to BALs, which utilise cells within the ECS. Unlike dialysis, which
provides intermittent filtration, in vitro studies of the BAK device
show differentiated active transport of essential nutrients, such as
glucose, amino acids and sodium bicarbonate, sufficient metabolic
activity and important endocrine processes [64]. Although the
current bioartificial systems have been utilised for clinical use,
there is scope as with the systems mentioned previously to use
them as a foundation for many other applications.
There have been two main groups focusing on the development
of a BAK device for clinical use. Humes et al. developed a device to
treat patients with acute renal failure and multiple organ failure
with a degree of success, whereas Saito et al. focused on the
prevention and treatment of long-term complications in patients
on maintenance dialysis [65]. However, both devices utilise the
same bioartificial component, hollow fibre modules seeded with
proximal tubule cells (PTCs), and pass ultrafiltered blood through
the module before administering back into the patient (Fig. 2).
Humes et al. [66] utilised commercially available haemofiltration
cartridges, with polysulfone (PSF)-based fibres seeded with human
PTCs (hPTCs). Although confluent monolayers of cells were
observed within the fibres [64], Saito et al., concurring with Oo
et al. [67], have commented on the lack of biocompatibility of PSF
and PSF-based membranes for generating a confluent monolayer
of cells [65]. The clinical trial undertaken with the BAK device from
Humes’ group has also been critically assessed, noting the lack of
documentation on expected effect size, the incompletion of treatment of patients assigned to continuous veno-venous haemofiltration(CVVH) + renal assist device (RAD) treatment and the lack
of statistical significance of the primary results with comparisons
performed as an as-treated rather than an as-randomised intention-to-treat sample [68].
Cell sources
To create an efficient, reliable model of the kidney in situ, a BAK
device must incorporate a cell type that can carry out several native
functions; therefore, research has been drawn to renal PTCs. These
cells perform a variety of renal-specific functions, including reabsorption, metabolic and transport functions, the secretion of
uraemic toxins and xenobiotics, and performing immunomodulatory functions [69]. Renal PTCs also have a range of transporters,
such as organic anion and organic cation transporters for basolateral drug uptake and ATP binding cassette (ABC) transporters
present on the apical surface for luminal excretion into urine.
Many of the transporters are polyspecific; that is, they accept
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compounds of different sizes and molecular structures, and have
overlapping substrate specificities [70].
Primary hPTCs are required for clinical BAK devices, but the sheer
number of cells needed (approximately 109 renal cells per device) is a
major obstacle to their widespread use, driving the search for alternative cell sources. Human PTCs have also been used as a wellcharacterised in vitro model of the kidney in drug transport studies,
owing to the classification of specific carrier proteins on the plasma
membrane by substrate specificity and functional assessment, and
the prediction of drug–drug interactions based on competitive
inhibition for transport [71]. Two transporters, organic anion transporter (OAT) 1 and OAT3, are thought to be the major OATs
responsible for the basolateral uptake of various organic ions from
the blood in vivo; OAT3 action was demonstrated by the mediation
of rosuvastatin uptake in vitro in the presence and absence of
probenecid, a uricosuric drug [72]. Probenecid has also been shown
to be a potent inhibitor of OAT2 and OAT4, where prostaglandin
F2a (PGF2a) uptake by OAT2, and estrone-3-sulfate uptake by apically located OAT4, were inhibited in the presence of probenecid.
The main uptake transporter of organic cations in hPTCs is thought
to be OCT2, located on the basolateral membrane. Functionality has
been demonstrated in hPTCs by the decrease in uptake of a fluorescent dye, ASP+ in the presence and absence of the inhibitor
quinine [73]. Although drug–drug interactions for organic cations
have been observed in other cell sources, such as LLC-PK1 cells, few
data on the drug interactions for OCT2 have been documented [74].
Similar to liver, P-gp is documented to transport a wide range of
xenobiotics in hPTCs, in particular anticancer agents, such as
methotrexate and cisplatin [75]. However, in hydrophilic statin
efflux, P-gp was found to have a minor role compared with the
MRP2 and breast cancer resistance protein (BCRP) efflux transporters [72]. Differences between effects of drugs in vivo and in vitro have
also been highlighted by the inhibitory effects of KW-3902 and
betamipron on OAT4 in vitro, but that showed no significant inhibitory effects on OAT4 in vivo [76].
The expression of both phase I [11 CYP enzymes and three
glutathione S-transferases (GSTs)] and phase II (three UGT and
variable expressions of three sulfotransferases) drug-metabolising
enzymes has also been found to be retained within primary hPTC
cultures. Furthermore, when grown in a confluent monolayer,
which is necessary in current BAK devices, the expression of these
enzymes was generally maintained at a measurable level, albeit
lower than in the native tissue [77]. Selective drug transporter loss
upon primary cell culture has been identified to be a consequence
of dedifferentiation. However, recent novel human tubular kidney
cell cultures grown on permeable filter supports have been shown
to express a wide range of drug transporters owing to the co-culture
of proximal and distal tubular cells [70]. This finding is concurrent
with the increased function from co-culture of hepatocytes and
nonparenchymal cells in BAL devices. Collectively, these results
indicate that primary cultures of hPTCs can metabolise drugs,
provide good models for drug interaction studies and benefit
the clinical setting when incorporated into BAK devices. Although
hPTCs do not always represent in vivo functionality, these results
could be improved by their incorporation into a dynamic bioreactor, providing a more accurate model of hPTCs in situ.
Using immortalised cell lines eliminates the problem of
needing vast numbers of cells for studies. Methods such as cloning,
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Drug Discovery Today Volume 18, Numbers 19/20 October 2013
derivation from cancerous cells and viral transfection can all produce genetically stable, but functionally altered cells. Human cells
are relatively difficult to immortalise and do not express differentiated features in vitro [64,78]. Therefore, the established cell line HK2, generated from the infection of human PTCs with an immortalising construct from the human papilloma virus, and the more
recently developed human kidney proximal tubular cell line
(HKC) from exposure of renal epithelial cells to Adeno 12:SV40
hybrid virus, do not provide an adequate replacement for primary
cells for use in clinical BAK devices [78]. Other established animalderived cell lines, such as LLC-RK1, LLC-PK1 and MDCKII, have also
provided adequate models for research, but interpretation to the
human situation is often difficult owing to species differences,
especially with regard to transporters and enzymes. Drug transport
with human renal cell lines has not been widely researched, possibly
owing to the substantial loss of differentiated function from the
immortalisation of human cells, and primary hPTCs or transfected
animal cell lines, such as MDCKII-MDR1, are more frequently used.
This highlights the challenge of generating a physiologically complex, well-characterised and validated in vitro cellular model that can
replace current US Food and Drug Administration (FDA)-recommended cell transport test systems, such as P-gp transfected MDCK
or LLC-PK1 cells [79].
Porcine PTCs are widely used in kidney research owing to their
reliability, and anatomical and physiological similarity to hPTCs.
In preclinical studies, porcine PTCs in a BAK device were shown to
tolerate the uraemic environment while providing metabolic,
reabsorptive and endocrinological activity in uraemic dogs [80],
enabling research to further investigate the benefits of BAKs, and
progress to hPTC utilisation for clinical trials. There are few studies
of the differences in drug interactions and metabolism between
porcine and human PTCs to date, although drug transporter
expression on primary porcine PTCs and cell lines are comparative
to native human kidney cells (Table 3). Both primary and porcine
cell lines are readily available and are an established model for drug
toxicity; however, an optimised model for drug toxicity would
require human PTCs instead of the readily available established
porcine model.
Owing to their multipotency and self-renewal capabilities,
embryonic stem cells (ESCs) might provide a reliable source of
PTCs for use in BAK devices when exposed to the right environmental cues [63]. Studies into human ESCs differentiating into
renal cells have not been widely undertaken; however, preliminary
studies have shown integration of undifferentiated and mesoderm-derived ESCs into the stroma and developing nephrons
[81]. They have also been shown to express a wide range of early
renal markers, and exhibit upregulation of kidney precursor markers, such as EYA1, LIM1 and CD24 [82]. However, the generation
of the different types of renal cell from human ESCs still needs to
be investigated thoroughly [81]. Owing to the sparse studies and
various obstacles, drug metabolism and transport by human ESCderived renal cells have not been documented, but will be essential
in any progression of stem cell utilisation towards potential in vitro
cell models used in pharmaceutical research.
Device designs
As mentioned above, the existing devices used for clinical
applications utilise haemofiltration cartridges seeded with hPTCs.
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TABLE 3
Transporter family
Gene product
Location
MDCK II
LLC-PK1
Primary
porcine cells
Refs
Organic anion transporters
OAT1 (SLC22A6)
Basolateral
X
X
[84–86]c
ABC transporters
Organic cation transporters
OAT2 (SLC22A7)
Basolateral
OAT3 (SLC22A8)
Basolateral
X
X
H
[84,85]
[84–86]c
OATP1A2 (SLCO1A2)
Apical
X
X
c,d
MRP1 (ABCC1); MRP5 (ABCC5)
Basolateral
H
H
H
[84–86]c,d
MRP2 (ABCC2); MRP4 (ABCC4)
Apical
H
H
H
[84–86]c,d
MRP3 (ABCC3)
Basolateral
H
[84,85]
MRP6 (ABCC6)
Basolateral
H
[84,85]
MDR1 (ABCB1)
Apical
H
H
H
[84–86]c,d
BCRP (ABCG2)
Apical
H
H
X
[84]
X
X
[84–86]
OCT1 (SLC22A1)
Basolateral
OCT2 (SLC22A2)
Basolateral
[84–86]
OCT3 (SLC22A3)
Basolateral
[84–86]
Carnitine and/or organic cation
transporters
OCTN1 (SLC22A4)
Apical
[84–86]
OCTN2 (SLC22A5)
Apical
[84–86]
Organic anion transporting polypeptides
OATP4C1 (SLCO4C1)
Basolateral
[86]d
Peptide transporters
PEPT1 (SLC5A1)
Apical
X
H
H
[84–86]c
PEPT2 (SLC5A2)
Apical
X
X
H
[84–86]
MATE1 (SLC47A1)
Apical
[86]
MATE2-K (SLC47A2)
Apical
[86]
Multidrug and toxic compound
extrusion transporters
a
Human drug transporter orthologues present on the cell lines and primary cells. Blank fields indicate no presence of transporters on the cells.
b
Abbreviations: H, present; X, absent; , only detectable after 35 cycles of PCR; , only present in freshly isolated samples.
c
Jin Jang, K., Ingber, D. (2011) Human kidney proximal tubule-on-a-chip for drug transporter studies and nephrotoxicity assessment. 15th International Conference on Miniaturized
Systems for Chemistry and Life Sciences, 2–6 October, 2011, Seattle, Washington, USA (http://www.rsc.org/images/LOC/2011/PDFs/Papers/502_0844.pdf ) [Accessed 21 May 2013].
d
FDA. Drug Interactions & Labeling: Drug Development and Drug Interactions: Table of Substrates, Inhibitors and Inducers. Available from: http://www.fda.gov/Drugs/
DevelopmentApprovalProcess/DevelopmentResources/DrugInteractionsLabeling/ucm093664.htm#major [Accessed 21 May 2013].
However, for an in vitro model, device design could be optimised.
For example, the number of cells needed could be greatly
decreased owing to the reduction in the size of the cartridge.
Hollow fibre cartridges could also be greatly optimised, because
commercially available modules have varying degrees of biocompatibility. Tailoring fibre materials to different cell types and
sources has been thought to improve cell performance and, for
primary hPTCs, it is thought that some of the membrane properties needed include a hydrophilic, negatively charged adhesive
surface [28]. When a variety of commercially used membranes,
such as regenerated cellulose (RC), PSF and/or PVP and polyethersulfone (PES) and/or polyvinylpyrrolidone (PVP) coated with
ECM, were tested with primary hPTCs, performance of the cells
was not improved, indicating that these materials are not suitable
for use in BAK devices. The differing properties between cell
sources are further highlighted by reports that PSF membranes
coated with ECM are able to maintain both primary porcine [83]
and LLC–PK1 cell monolayers for 3 weeks after confluency [84].
Coating commercially available fibres with tailored ECM is
preferable to tailoring membrane materials owing to the ease of
assembly and cost effectiveness. Various single and double coatings have been tested, including collagen IV and laminin blends
with a variety of growth factors [69]. Single coatings demonstrated
little improvement on hPTC performance within the fibre,
whereas double coatings of 3,4-dihydroxy-L-phenylalanine
(DOPA) and collagen IV on haemocompatible membranes
revealed functional epithelia formed by hPTCs in a HFB [67].
Bioreactor configuration should also be taken into consideration
to both achieve native mimicry and produce an optimal environment for the cells to proliferate. Previous bioartificial devices
utilising hollow fibre cartridges have encountered problems with
the nonuniform distribution of cells and the lack of transport of
gases and nutrients [85]; therefore, including an oxygenator or
aerated medium would need to be considered. Additionally, to
mimic the kidney, the device would need to include a tubule
(hollow fibre) and capillary and/or blood (extracapillary space)
section with shear rates according to native conditions. These
would be higher in the ECS (typically 10 dyn/cm2 on endothelial
cells in vivo) than in the fibre (typically 1 dyn/cm2 on PTCs in vivo)
[86]. This could lead to movement of molecules through the
membrane if the monolayer of cells was not tight, and affect cell
performance owing to varying pressures. However, from previous
devices used clincally, no loss of membrane integrity has been
reported when using different flow rates in the tubule and ECS
[66].
Emerging in vitro models have also adhered to the established
hollow fibre structure. Researchers at the Wyss Institute, Harvard
University have introduced ‘organs on a chip’. These microfluidic
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Drug transporter orthologues present on MDCKII and LLC-PK1 cell lines and primary porcine PTCsa,b
Drug Discovery Today Volume 18, Numbers 19/20 October 2013
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PDMS channel
Porous
membrane
PDMS reservoir
Reviews FOUNDATION REVIEW
35-mm culture dish
Kidney tubular cells
apposed to porous membrane
Drug Discovery Today
FIG. 3
Schematic of human kidney proximal tubule-on-a-chip in the form of a multilayer microfluidic device that integrates a polydimethylsiloxane (PDMS) microfluidic
channel, a porous membrane and a PDMS reservoir. Human primary proximal tubule cells were cultured on the device.
Modified, with permission, from Jin Jang et al.4
devices, in this case ‘kidney on a chip’, utilise hPTCs seeded within
a microfluidic channel apposed to a polyester membrane. The
channel is immersed in medium, and a shear stress of 0.2 dyn/cm2
is applied through the channel (Fig. 3). Cytoskeletal proteins, tight
junction proteins, ion transporters and specific drug transporters
(e.g. OCT2) have shown improved results in a fluidic model
compared with a static version, and nephrotoxic compounds,
such as cisplatin, have shown the ability of the microsystem to
recapitulate the in vivo toxicity of the drugs. Overall, the system
presents a substantial answer to the need of a reliable 3D in vitro
model.4 However, functional characterisation of the drug transporters within the device that are acknowledged to be of clinical
importance to drug disposition by the FDA, namely P-gp, BCRP,
OCT2, OAT1 and OAT3, should be undertaken to establish an ideal
model system.5
Concluding remarks and future perspectives
Bioreactors have been utilised in many different areas of industry
and research and, for the progression of in vitro models of 2D
cultures and suspensions to 3D constructs mimicking the natural
physiological state in situ, they have proven to be crucial. The
generation of bioartificial devices which provide both the structural and biological aspects, needed for cell growth and function,
have been exploited for clinical and research purposes. The most
established bioartificial system, the bioartificial liver, is a prime
example of the development of a bioreactor designed initially for
therapeutic use, being utilised as an in vitro model. From simple
perfused bed and HFBs, focus on the native liver structure and
4
Jin Jang, K. and Ingber, D. (2011) Human kidney proximal tubule-on-a-chip for
drug transporter studies and nephrotoxicity assessment. 15th International
Conference on Miniaturized Systems for Chemistry and Life Sciences, 2–6
October, 2011, Seattle, Washington, USA (http://www.rsc.org/images/LOC/
2011/PDFs/Papers/502_0844.pdf ) [Accessed 21 May 2013].
5
FDA. Drug Interactions & Labeling: Drug Development and Drug Interactions: Table of Substrates, Inhibitors and Inducers. Available from: http://
www.fda.gov/Drugs/DevelopmentApprovalProcess/DevelopmentResources/
DrugInteractionsLabeling/ucm093664.htm#major [Accessed 21 May 2013].
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cellular functionality within the device has prompted multifibre
bundle designs.
In addition to bioreactor design, development of the BBB models has highlighted the importance of cell sources, especially with
co-culture of different cell types, within a fluidic system opposed
to a static culture. It is understood that primary human cells are
the gold standard for an in vitro model; however, owing to the
difficulty of acquisition, BBB models have had to utilise other
primary sources. Embryonic stem cells could be a possible solution
to the dilemma of availability for the future; however, more
research is needed to investigate the difference in properties
between primitive and mature human CNS and ECs before incorporation into a device for DMPK investigation.
Bioartificial kidney devices have followed the progression of
BAL systems. From development for clinical use as a renal replacement therapy for acute and chronic renal conditions, they are now
starting to reveal their potential as in vitro models (Fig. 4). Primary
human cells have been utilised within the system, and are relatively easy to acquire and maintain, especially with the reduced
cell number requirement in a model. However, co-culture with
distal tubule cells has not been fully investigated within a BAK
device, and functional characterisation of drug transporters crucial
for ADMET studies have not yet been undertaken in the existing
microfluidic system.
To date, there are no universal criteria laid out to aid the
development of an acceptable and/or suitable bioreactor. However, it is widely accepted that in vivo micro environmental
aspects, such as cell viability, cell–cell interactions, cell–matrix
interactions, tissue architecture and cell oxygenation, are vital to
give an accurate representation of the physiological state and,
consequently, the effects of new chemical entities [3]. Therefore,
it is important to bear in mind those factors that are specific to the
desired tissue when developing in vitro models for use in pharmaceutical research. Furthermore, it is noticeable that little or no
measurements are currently taken from these bioreactor-based
devices to help control or monitor cell health or functionality
during use. Thus, we anticipate that, with the continuing
advances being made in sensor technology, the next generation
Drug Discovery Today Volume 18, Numbers 19/20 October 2013
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Functional cell type
Microenvironment
Progression
of
bioreactors
as in vitro
models
Cell type and sources
Primary cells, cell lines
Cell characterisation
Membrane material and
properties
Transporter,expression, metabolic enzyme
possession, native morphology, adherent
protein expression, cell functionality
Material, pore size, porosity, surface
charge, hydrophilicity, topography,
shear rate
Functional cell support
Coatings
Similarity to native
microenvironment, number of
coatings, mixture of coatings
Co-culture of cells
Bioartificial Bioartificial
kidney
BBB
Bioartificial
liver
Future developments- Utilising stem cell
derived specialised cells
Primary human hepatocytes
Transporters:MRP2, MDR1, BCRP
Metabolic enzymes: CYP1A, CYP2C9
CYP3A
Cell functionality markers: Albumin
synthesis, lactate dehydrogenase,
aspartate aminotransferase, glucose
secretion, lactate production
Co-culture with non-parenchymal cells
PBMEC/C1 (porcine brain
microvascular endothelial cells), C6 (rat
glioma cells)
Glucose consumption, lactate
production, paracellular transport,
benzodiazapene permeation
Human proximal tubule cells
γ-glutamyl transferase activity,
Future developments- Tools for
real time monitoring of cells
incorporated e.g. gas/ electrolyte
and metabolite sensors
Multicompartment hollow fibre
bioreactor
Native architecture
Device type
Hollow fibre and perfusion systems
Device setup
Single circuit, multi-compartmental
devices
Future developments- Miniaturisation
of devices
Polyethersulfone,
hydrophobic
multilaminate
hollow fibres
Commercially available hollow
fibre cartridge
Polypropylene, 0.5 µm
pores, 4-50 ml/min
Fibronectin
Single fibre hollow fibre
bioreactor
Polyethersulfone/
polyvinylpyrrolidone,
0.5 µm pores, 80 µl/min
Double coating of DOPA
and human collagen IV
parathyroid hormone, RT-PCR, SEM,
immunostaining
Drug Discovery Today
FIG. 4
Key aspects from the progression of bioreactor-based in vitro models and future developments [21,60,67]. Abbreviations: BBB, blood–brain barrier; BRCP, breast
cancer resistance protein; CYP, cytochrome P450; DOPA, 3,4-dihydroxy-L-phenylalanine; MDR1, multidrug resistance protein; MRP2, multidrug resistanceassociated protein.
of devices will incorporate these to improve either device function or longevity.
Bioartificial systems have started to show great promise in the
generation of in vitro models for various organs and disease states.
Observations from existing systems are the importance of mimicry
of both structure and environments in situ in the bioreactor, the
choice of cell and source to model the functions of the organ and
essential factors for maintaining cellular function, such as oxygen
availability and nutrient and metabolite delivery. With these
factors in mind, successful models utilising bioreactors can be
achieved in the near future and will hopefully provide a better
means to predict clinical outcome from in vitro data.
References
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2 Scherer, F. (2011) R&D Costs and Productivity in Biopharmaceuticals. F Kennedy School
of Government, Harvard University
3 Astashkina, A. et al. (2012) A critical evaluation of in vitro cell culture models for
high-throughput drug screening and toxicity. Pharmacol. Ther. 134, 82–106
4 Ballard, P. et al. (2012) The right compound in the right assay at the right time: an
integrated discovery DMPK strategy. Drug Metab. Rev. 44, 224–252
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